The invention relates to nuclear medicine, and more particularly relates to scintillation camera systems for use in nuclear medicine studies. Scintillation cameras are used to image the distribution of gamma ray radioactive material within a body part or organ, such as the brain or the breast, for example, for diagnostic purposes. A source of penetrating radiation is administered to the patient, which typically consists of a pharmaceutical tagged with a gamma-ray emitting radiopharmaceutical designed to go to and deposit in the organ or elements of the body under diagnostic examination, such as, in the detection of a tumor. Gamma rays emitted by the radiopharmaceutical are received and detected by the camera, the position of each detected ray event is determined, and the image of the radioactivity distribution in the organ or other body part is constructed by known techniques from an accumulation of events.
Scintillation cameras generally employ one or more optically continuous crystals of thallium activated sodium iodide, Nal (TI), as the gamma ray energy transducer. Energy of the gamma rays are absorbed in the crystal and are converted to light emissions called scintillation events, each event having an energy proportional to the energy of the absorbed gamma ray. In conventional cameras, light is transmitted from the crystal to an optically clear glass window through a silicone gel interface that fills a thin separation between the glass window and the crystal. The optical window is part of a container that seals the crystal from air and humidity, which would otherwise oxidize the crystal and degrade its optical clarity. An array of photomultiplier tubes (PMT) is optically coupled to the glass window, typically by means of optically coupling grease, in order to transmit light to photocathodes located on the inner surface of the glass entrance face of each photomultiplier tube. Thus, the scintillation light events must pass sequentially from the Nal (TI) crystal through the silicone gel interface, glass window, silicone grease interface, and photomultiplier glass before striking the photocathodes within the photomultiplier tubes. The photocathodes serve to convert the light to electrons by the photoelectric effect and the electrons are multiplied in the photomultiplier tubes.
Amplified signals generated in photomultiplier tubes in the vicinity of the scintillation event are then mathematically combined by analog or digital means in an attempt to determine the position and the energy of the gamma-ray absorption in the crystal. Accurate determination of the energy level and position of the scintillation event requires that the efficiency of transmission of the scintillation light to the photomultiplier tubes be high. Light dispersion or deflection adversely modifies the ideal light distribution and degrades position determination. If light is reflected back from an interface or undergoes multiple reflections before striking a photocathode, the position information contained in the photomultiplier signals is likely to be compromised. Accurate position determination of scintillation events is essential for high quality resolution.
Because of the number of intervening sources that degrade accurate positioning of scintillation events, such as those listed above, there is a need for an improved method to accurately localize scintillation events and thus improve resolution.
The present innovative concept provides a means to improve the realized energy resolution of a scintillator based ionizing photon detector by using information of the Locus of the Scintillation (LOS) photons resulting from an absorption within the scintillator of an ionizing photon. The LOS is then used to select ionizing photon interaction events on the bases of the location of their interaction in the scintillator.
A scintillator crystal is mounted to a photomultiplier tube, which is acting as the primary detector. One of several methods is then employed to determine the LOS resulting from the absorption of the gamma ray. The primary detector can be a position sensitive photomultiplier tube (PS-PMT) with crossed wire anodes whose outputs are read simultaneously for each scintillation event, for example. In the last ten y ears, reliable high performance compact position sensitive photomultiplier tubes, such as the Hamamatsu R2486 and the Hamamatsu R3292 have been made available from industry for research.
The broadness of the distribution of the electron cloud on the anode wires is dependent upon the LOS light resulting from the x-rays or gamma rays. Ionizing photons that interact towards the top of the scintillating crystal will produce lower signals, due to absorption in the scintillator, than photons that interact closer to the PMT surface. These photons that interact toward the top of the scintillator crystal give rise to a broader electron cloud distribution on the photocathode which in turn results in a broader distribution at the anode outputs. Conversely, photons with higher energy travel further through the scintillation crystal and interact closer to the PS-PMT. These higher energy photons result in a narrower electron cloud distribution at the anode outputs. By selecting scintillation events based on the size of the electron cloud distribution at the anode outputs one can obtain enhanced energy resolution for the detector by correcting for the depth of interaction dependence.
This method is very useful in nuclear medicine gamma cameras by improving the rejection of scattered radiation that would otherwise contaminate the image. No photon statistics would be lost since all of the interaction events would still be acquired. The only difference in using the present method is that the events would be sorted to different image displays for different energy ranges of the detected gamma rays. This would provide several options to the radiologist reading the image. The clinician would then have the advantage of looking at all the image data as usual and could have specific energy windows enhanced by making use of the depth of interaction information.